Positron emission tomography (PET) is a method of imaging that uses radiolabeled probe molecules to target, detect, and quantify biological processes in vivo. PET techniques are used to study disease mechanisms, to develop new diagnostic and therapeutic methods, to detect early stage disease, and to monitor responses to therapies. The equipment, infrastructure, and personnel currently required to produce PET probes severely constrain the availability and diversity of probes, hindering advances in disease diagnosis, therapy, and medical research that requires this imaging method.
The approach to synthesis of biochemical compounds with radioactive nuclei generally starts with a charged particle accelerator. Particle accelerators have the following attributes: an ion source, electrostatic extraction optics that select a single polarity of ion for acceleration, electromagnetic fields to accelerate and focus the ions, a vacuum chamber to prevent elastic and inelastic scattering of the ion beam, collimation apertures, and external shielding to protect operators and electronics from neutron and ionizing radiation produced in the accelerator. Referring to FIG. 1, positive or negative ions are formed in an ion source (101), typically by electron impact, then separated by polarity (anions from cations) and mass (atomic ions from molecular ions and electrons) and accelerated in a linear or cyclotron accelerator (102) with electromagnetic fields to increase their kinetic energy. The charged beam is then extracted from the accelerator (103), collimated, and shaped using electrostatic lenses. Approximately 20% of the beam current is lost in the cyclotron, contaminating the housing with heavy radioactive nuclei and neutron radiation. Extraction of negative ions such as 1H− is also accomplished with electrostatic fields. These anions must then be converted to protons by passing them through a carbon foil to strip two electrons with almost 100% efficiency. Since energetic negative ions do not undergo nuclear reactions with the metallic accelerator components and the reactive positive ion beam has a short path, activation of the housing is reduced. However, the acceleration of negative ions requires ultra high vacuum (<10−10 atmospheres) to mitigate charge neutralization. The acceleration of all charged species also produces electromagnetic radiation; at the energies required for subsequent formation of radionuclides, this is ionizing radiation that requires heavy, bulky shielding (104) for safe operation.
Effective formation and acceleration of ions by electromagnetic fields requires operation in vacuum chamber (105), so the next step impinges the energetic ion flux through a window material (105) and onto a solid, liquid, or gas precursor material (106). The energetic ions convert some of the precursor (106) to radionuclides (107) by nuclear reactions. The mixture of precursor and radionuclides (106, 107) are transferred (108) to a separately shielded (109) hot cell or microfluidic reactor where chemical reactions (110) and purifications (111) convert the radionuclide into an injectable radiochemical reagent.
The collision of the accelerated ions with the precursor material occasionally results in a nuclear reaction whose probability is quantified as the integral of the product of a cross-section Q(E), the energy distribution of the ion flux (f(E)), and the relative velocity of the ion and precursor nuclei (v(E)). The rate of radionuclide (RN) production from a concentration of precursor is given by
                                          ⅆ                          [              RN              ]                                            ⅆ            t                          =                              [            Precursor            ]                    *                      ∫                                          Q                ⁡                                  (                  E                  )                                            *                              f                ⁡                                  (                  E                  )                                            *                              v                ⁡                                  (                  E                  )                                            ⁢                              ⅆ                E                                                                        (                  Eq          ⁢                                          ⁢          1                )            
These nuclear reactions yield an unstable material that decays by releasing a positron, which in turn collides with an ambient electron to produce two counter-propagating gamma rays. The gamma rays are then recorded by coincidence detection in a toroidal sensor. Following tomographic inversion the location of the decaying radionuclide can be determined to within fractions of a millimeter. PET imaging has been applied to the diagnosis of vascular function (Laking et al., The British Journal of Radiology, 76 (2003), S50-S59 E), arthritis (Bruijnen et al., Arthritis Care & Research Vol. 66, No. 1, January 2014, pp 120-130), and tumerogenesis (Aluaddin, Am J Nucl Med Mol Imaging 2012;2(1):55-76), among many others.
The specific activity (SA) of a radioactive tracer is an important figure of merit for a PET reagent. It is defined as the intensity of radiation divided by the mass or number of moles of material, and it decreases with time (t) according to the expression exp(−t/τ) where the decay rate (1/τ) is a fundamental property of the specific radionuclide. This decay begins the moment a radionuclide is formed, and extensive research has been devoted to methods of swiftly and efficiently inserting the radionuclide into a biological probe through chemical reactions and purifications to produce a PET reagent in the shortest possible times.
Representative values of τ are listed in table I. Small values of τ imply rapid decay, which is advantageous because it produces more decay events per second and therefore greater signal to noise ratios when collecting image data. However, for these same values of τ any factor that increases t leads to a faster loss of potency of the reagent.
TABLE IProperties of four representative medical isotopes that are produced by proton bombardment.MedicalDecay time (τ)NuclearEnergyYieldIsotopeminutesReaction(MeV)(milliCi @ sat)11C29.311B (p, n)8-20 40/μA11C29.314N (p, α)12100/μA13N14.413C (p, n)5-10115/μA13N14.416O (p, α)8-18 65/μA15O2.9415N (p, n)4-10 47/μA15O2.9416O (p, pn)>26 25/μA18F15818O (p, n)8-17180/μA
One problem with the current methods is their requirement for an accelerator or cyclotron to produce the ion beam from which radionuclides are formed. Cyclotrons require heavy and expensive magnets, high voltages, substantial electric power, and extensive radiation shielding. For example, Bhaskar Mukherjee has summarized the shielding requirements in Optimisation of the Radiation Shielding of Medical Cyclotrons using a Genetic Algorithm, which is incorporated herein by reference in its entirety. According to Mukherjee, “[t]he important radioisotopes produced by Medical Cyclotrons for present day diagnostic nuclear medicine include 201T1 (T1/2=73.06 h) and 67Ga (T1/2=78.26 h). These radioisotopes are generated by bombarding the thick copper substrates electroplated with enriched parent target materials with 30 MeV protons at ˜400 μA beam current. The target bombardments result in the production of intense fields of high-energy neutrons and gamma rays.” A summary of medical cyclotron characteristics abstracted from a presentation by Jean-Marie Le Goff, [A very low energy cyclotron for PET isotope Production, European Physical Society Technology and Innovation Workshop Erice, 22-24 Oct. 2012] is reproduced in Table II. As can be seen with reference to Table II, the average weight of a medical cyclotron is 36 tons, the average weight required for shielding is 47 tons, and the average power requirement is 101 kilowatts. The smallest device in Table II has a total weight of ten tons and requires 10 kW of power. In other words, the size, weight, and power of a cyclotron require that it be placed in a fixed installation.
TABLE IIParameters including size, weight, and power of some commercial cyclotrons thatare used for medical isotope production.BeamCycltronShieldCompanyCyclotronEao:rgyCurrentIonRF Frequ.WeightWeightPowerNameModelParticles(MeV)(μA)Source(MHz)(tons)(tons)(kW)ACSITR14H−14>100 Cusp74224060ACSITR19(9)H−(D−)19, 9 >300 (100)Cusp74 (37)2265ACSITR24H−24>300 Cusp83.58480ACSITR30(15)H−, (D−)30, 151500 (400)Cusp56150ABTTabletopH+  7.5 5PIG723.27.610BestBSCI 14pH•14100PIG731460BestBSCI 35pH−15-351500 Cusp7055280BestBSCI 70pH−70800Cusp58195400CIAECYCCIAE14H−14400CuspCIAECYCCJAE70H−70750CuspNIIEFACC-18/9H−, (D−)18, 9 100 (50)Cusp38.220Feb-00EUROMEVIsotraceH•12100Cusp1083.840GEMINItraceH−  9.6>50PIG10194035GEPETtraceH−, (D−)16.5, 18.6 >100 (6.5)PIG27224770IBACyclone 3D+  3.8 60PIG14514IBACyclone10/5H−, (D−)10, 5 >100 (35) PIG42124035IBACyclone11H+11120PIG42135235IBACyclone18/9H−, (D−)18, 9 150 (40)PIG422550IBACyclone30H−, (D−)30, 151500 Cusp50180H−, (D−)350 (50)(50)IBACyclone70H2+, He++30-70, 15-5  (35)66 (30)125350KIRAMSKIRAMS-30H•15-30500Cusp64KIRAMSKotrun-13H+40100PIG77.32080187SiemensEclipseRDH−112 × 40PIG113935SiemensEclipseHI/STH−112 × 40PIG7235SumitomoHM-7H−, (D−)7.5, 3.830SumitomoHM-10H−, (D−)9.6, 4.852SumitomoHM-12/5H−, (D−)12, 6  60 (30)PIG45115645SumitomoHM-18H−, (D−)18, 10 90 (50)PIG45248655Average3647101
A second problem with PET isotope synthesis stems from the fact that materials prepared at the fixed cyclotron site lose specific activity during transport to the site where patients are scanned. This problem is particularly acute when the transport time t_transport is long compared to the decay time τ, because the specific activity drops by exp(−ttransport/τ).
A third problem results from the economics of producing the reagents at a central site. In order to spread the capital and operating costs of the facility many doses must be made at once, and these must be distributed in a timely manner to patients at dispersed locations. This complicates the logistics of patient care because scanning facilities must be choreographed with the production schedule of the cyclotron while accounting for material degradation in transit.
Yet another problem is that isotopes with very short lifetimes (small values of τ) cannot be used except in very close proximity to the accelerator because their specific activity degrades too rapidly to permit detection with useful signal to noise ratios in a PET scanner. For example, the half-life of H215O, a PET tracer used to measure perfusion in cardiac imaging, is only 2 minutes.
Another problem is that production of multiple doses at once requires higher beam currents, which in turn demand windows between the vacuum and precursor regions that can manage thermo-mechanical stresses without significantly degrading the energy or current of the ion beam. A second problem with higher beam currents is collateral radiation damage to the chemical composition of the precursor. The irradiation of a large protein molecule containing nitrogen with large currents of 2H+ ions from a cyclotron to synthesize 15O radiolabels, for example, may degrade or denature the protein. This collateral damage limits the range of precursor materials to those that resist radiation damage, such as H218O, one precursor for production of 18F by proton beams.
Once a radionuclide is formed it can be chemically bound into a molecule that serves to mark specific molecular or biological activity. For example, 18F is produced from H218O as aqueous 18F− anions that are converted through one or more chemical reactions to 18F-fluoro-deoxyglucose. This injectable reagent is taken up in vivo by cells and accumulates in their mitochondria, providing an indication of cellular metabolism rates. These chemical reactions and purifications are performed in heavily shielded enclosures or ‘hot cells’, named so due to the large amount of shielding required to prevent radiation exposure to the operators. The typical reaction volume of “hot cells” is of the order of 1 milliliter (mL) though the amount of radioactive atoms or molecules present is extremely small, typically 6×1011 atoms or molecules. A typical processing time processing (tprocess) is 40-50 minutes, that with the exception of 18F, exceeding by far the decay time of most interesting RN. The time and care required for this manual conversion contributes significantly to loss of specific activity in the final product.
Van Dam et al. disclosed a significant improvement in U.S. Pat. No. 7,829,032, entitled Fully Automated Microfluidic System for the Synthesis of Radiolabeled Biomarkers for Positron Emission Tomography, which is incorporated herein by reference in its entirety. Incorporating small-volume, automated processing substantially reduced the time required to convert radioactive precursors to injectable reagents, enabling higher specific activity and safer production than prior methods. However, a limitation of this approach is that it separates production of the radioisotope from chemical conversion, so the time to transfer radionuclides between a cyclotron and the microfluidic system (ttransfer), indicated schematically by (108) in FIG. 1, contributes to loss of specific activity according to equation (1).
U.S. Pat. No. 8,080,815 discloses use of microfluidic systems to synthesize radioactive tracers, which is incorporated herein by reference in its entirety. This reference discloses use of commercial micro-fluidic technology to process radionuclides created by a small cyclotron accelerator that separately produces radionuclide for one dose for human image needs, for example approximately 10 milliCurie (mCi) for 18F-fluoro-deoxyglucose. This method suffers from all of the shielding and auxiliary deficiencies of electromagnetic accelerators, and also from the need to convey radionuclides from the cyclotron to the microfluidic reactor as indicated by (108) in FIG. 1.
Referring to FIGS. 3 and 4, charged particle accelerators have the following attributes: (1) an ion source system, (2) magnetic and/or electric fields that form and accelerate beams of single polarity charged particles with energy sufficient to undergo nuclear reactions, (3) a target for irradiation by the charged particle beams, and (4) a shielding system. Cyclotron accelerators were introduced in 1932 by E. O. Lawrence, who received the 1939 Nobel Prize for “the invention and development of the cyclotron and for results obtained with it, especially with regard to artificial radioactive elements.” Cyclotrons and linear accelerators require a stream of particles of only one polarity because they use a combination of fixed and oscillatory electromagnetic fields that produce opposite forces on charges of different polarity. These beams are streams of particles whose center of mass moves with high velocity while its spread in energy, ΔE, is smaller than its energy E (ΔE/E<<1). Note that as ΔE approaches E the divergence of the beam increases, obviating further acceleration and directing toward targets. Cyclotrons have been widely used for production of radioisotopes and are commercially available, as summarized in Table II. However, the acceleration of the charged particles generates electromagnetic radiation that can damage electronics and is hazardous to human operators. These large, complex machines require kilowatts of electric power and many tons of radiation shielding. Moreover, the use of high voltages in vacuum requires careful shielding and insulation, contributing to the complexity and expense of conventional accelerators.
Efficient generation of radionuclides requires maximizing the integrated product of the velocity-weighted energy distribution f(E)*v(E) with the cross section Q(E) in equation 1 above. Another problem with accelerator-based radionuclide synthesis is that the resulting ion beams generally have energies well above that for which the radionuclide precursor has its maximum cross section. This in turn requires larger currents to increase the production rate, concurrently increasing collateral radiation damage to the precursor materials.
Accordingly, there exists a need for additional devices and methods for production of radioactive reagents, and in particular, devices and methods that avoid the aforementioned limitations. Such devices and methods would be particularly useful in nuclear medicine, including positron emission tomography.